Biological assays (commonly referred to as bioassays) are a type of scientific experiment. Bioassays are typically conducted to measure the effects of a substance on a living organism and exploit in many instances the concentration, purity or biological activity of a substance such as vitamin, hormone, and plant growth factor. While measuring the effect on an organism, tissue cells, enzymes or receptor bioassays may also process and/or compare a standard preparation, i.e. a reference. Bioassays may be qualitative or quantitative and are typically employed in healthcare, environmental analysis, and emergency/threat type applications.
Healthcare: where bioassays may provide personalised access for individuals, relatives, carers and other specialists to real-time or historical information generated by wearable sensors, implantable devices or home based diagnostics units facilitates home or community based healthcare as well as improving diagnosis and patient data acquisition for hospital-centric treatments. Further, access to low cost communications and diagnostics also provides a means to rapidly improve the delivery of healthcare in less well-developed regions and remote regions.
Environment: where bioassays may provide multi-chemical sensors monitoring air and water quality can provide early warning of pollution events arising at industrial plants, landfill sites, reservoirs, and water distribution systems at remote locations for example. Increased analytics and wider/denser sensor networks may provide enhanced detection/categorization of events as they happen, and organizing/prioritizing the response(s).
Emergency/Disaster and Threat Detection: wherein bioassays may provide multi-measurand analytics and wider/denser/faster/immediate analytics against chemical, biological or radiological threats. Bioassays may therefore adjust chemical and biological measurements which today are overwhelmingly post-event and primarily related to gathering remedial and forensic information.
Microbiological testing, for example, demonstrates how microfluidics may provide benefit in one respect wherein traditionally, the plating and culturing to determine cells counts of bacteria requires multiple biochemical and serological characterization steps typically requiring days to weeks and hence unsuitable for many applications. Accordingly, developments of alternatives including emerging technologies such as enzyme linked immunosorbent assay (ELISA), polymerase chain reaction (PCR), DNA and flow cytometry have been geared to increasing speed of detection and reducing the volume of sample required. Pathogen detection utilizing ELISA has also become well established. However, integrated microfluidic systems (also referred to as “lab-on-a-chip”) offer improvements in the mass transport of the bacteria to the sensors and reductions in detection time to below 30 minutes.
To date a major challenge for the clinical use of “lab-on-a-chip” (LOAC) systems has been that they are generally complex and require sophisticated peripheral equipment, and as a result have proven much more difficult than anticipated to implement as low cost, robust and portable point-of-care systems. Whilst less integrated solutions have also been developed these are generally categorized as biosensors. Biosensors basically incorporate biological recognition elements (probes) such as antibodies, nucleic acids, and other types of receptors which provide a specific affinity toward a target analyte, and a transducer that converts the ensuing recognition event and biochemical activity into a measurable signal (commonly optical or electrical in nature). A wide range of biosensors has been developed for selective bacteria detection. However, most of these sensors rely on antibodies for capturing the bacteria on the surface, but antibodies are not efficient at capturing bacteria (bacteria are enormous compared to small molecules against which antibodies are commonly used), In addition, to date only a few biosensors may be integrated into LOAC systems in order to allow them to produce inexpensive LOAC systems which are easily amenable to miniaturization and mass production, and are potentially portable, using compact possibly hand-held instruments, using reusable or disposable detectors.
Another class of sensors that has progressed tremendously over the last few years are circulating tumor cell sensors. Circulating tumor cells (CTCs) are shed by tumors into the blood and are key to metastasis and an important prognostic for cancer progression. However, the isolation of CTCs is extremely challenging as they are significant even at ultralow concentrations of 5 cells per milliliter and the difficulty of separating them from white blood cells that are of similar size but present at concentrations of several million times higher, approximately 107 cells per milliliter. Further, CTCs must be selected against a background of immune cells that are similarly a million times more concentrated. Such imbalances in biological material to be detected versus the background biological environment match or exceed those with bacteria such as methicillin-resistant staphylococcus aureus (MRSA) which is a bacterium responsible for several difficult-to-treat infections in humans.
Microfluidics provided a breakthrough in CTC cell isolation, see for example Nagrath et at in “Isolation of Rare Circulating Tumour Cells in Cancer Patients by Microchip Technology” (Nature, 2007, pp. 1235-1239), Adams et al in “Highly Efficient Circulating Tumor Cell Isolation from Whole Blood and Label-Free Enumeration Using Polymer-Based Microfluidics with an Integrated Conductivity Sensor” (J Am. Chem. Soc., Vol. 130(27), pp 8633-8641), and S. Stott et at in “Isolation of Circulating Tumor Cells using a Microvortex-Generating Herringbone-Chip” (Proc. Nat. Acad. Sci., Vol. 107(43), pp 18392-18397). However, a drawback of these type of microfluidic technologies developed to date was that their low flow rate typically resulted in isolation times of several hours. In contrast D. Juneker in U.S. Patent Application 2012/0,617,714 entitled “Methods and Devices for Multi-Dimensional Separation, Isolation and Characterization of Circulating Tumour Cells,” the entire contents of which are included by reference, overcomes this drawback with a large conduit based sequential filtering platform which can be adapted to microfluidic based LOAC systems such as described below in respect of embodiments of the invention.
Referring to FIG. 1A the operating principle of microfluidics is outlined. For a liquid on a surface the surface tensions of the liquid, γSL, γLG, γGS being the solid-liquid, liquid-gas, and gas-solid defines a contact angleθC as given by Equation (1A) below. Re-arranging Equation (1A) yields Equation (1B) that defines the contact angle as a function of the surface tensions of the liquid, γSL, γLG, and γGS respectively. When 90°<θ<180° the liquid is non-wetting on the surface and when 0°<θ<90° the liquid wets the solid surface. When θ=0° a special condition known as wetting out is achieved, which implies γSG=γSL. Accordingly wetting of a liquid can be promoted by a variety of techniques including, but not limited to, roughening the surface (if θC<90°), reducing surface tension through addition of a surfactant to the liquid, putting an adsorbate material on the solid, and chemically modifying the solid surface (e.g. plasma treatment). Poly(dimethylsiloxane) (PDMS) which is one of several types of silicones employed in a variety of medical devices as well as contact lenses has θwater≈105° in its untreated form such that water does not wet. However, immediately after plasma treatment of PDMS θC=0°.
The Young-Laplace equation, see Equation (2), describes the capillary pressure, PC, across the interface between two static fluids due to the phenomenon of surface tension and states that this pressure difference is proportional to the surface tension, γ, and inversely proportional to the effective radius, r, of the interface and also depends on the wetting angle, θ, of the liquid on the surface of the capillary as given by Equation (2A) below. As evident in FIG. 1B the situation for a film is more complicated as a liquid film of dimensions L, w, and d in the x, y, and z dimensions respectively with exhibit a first radius r1 in the y direction and r2 in the x direction. Solving the various equations yields Equation (2B) wherein it is evident that the pressure differential, ΔPC, can therefore be adjusted through variations in the dimensions of the microfluidic channel within which the fluid is constrained, where θb,t,l,r are the contact angles of the liquid with the bottom, top, left and right walls of a microfluidic channel respectively.
                                          γ            SL                    -                      γ            SG                    -                                    γ              LG                        ⁢            cos            ⁢                                                  ⁢                          θ              C                                      =        0                            (                  1          ⁢          A                )                                          cos          ⁢                                          ⁢                      θ            C                          =                                            γ              SG                        -                          γ              SL                                            γ            LG                                              (                  1          ⁢          B                )                                          P          C                =                              2            ⁢            γ            ⁢                                                  ⁢            cos            ⁢                                                  ⁢            θ                    r                                    (                  2          ⁢          A                )                                          Δ          ⁢                                          ⁢                      P            C                          =                  -                                    γ              ⁡                              (                                                                                                    cos                        ⁢                                                                                                  ⁢                                                  θ                          b                                                                    -                                              cos                        ⁢                                                                                                  ⁢                                                  θ                          t                                                                                      d                                    +                                                                                    cos                        ⁢                                                                                                  ⁢                                                  θ                          l                                                                    -                                              cos                        ⁢                                                                                                  ⁢                                                  θ                          r                                                                                      w                                                  )                                      .                                              (                  2          ⁢          B                )            
To date multiple microfluidic systems have been developed targeting so called “point-of-care” (POC) analysis but with very few exceptions these have to date relied upon complex peripherals for system operation, and sometimes for assay readout as well. This has limited their adoption within clinics to date as well as preventing their deployment in more consumer driven POC analysis outside clinics without medical supervision integrating with online electronic health records, see for example Steinbrook in “Personally Controlled Online Health Data—The Next Big Thing in Medical Care” (New England J. Med., 2008, pp. 1653-1656).
FIG. 2A depicts an example of a conceptual microfluidic POC device after Gervais et al in “Microfluidic Chips for Point-of-Care Immunodiagnostics” (Advanced Materials, Vol. 23(24), pp. H151-H176, hereinafter Gervais1) wherein a POC tester comprising a body 210A and cover 210B allows a user to perform a measurement or measurements based upon the provisioning of a sample and its initial processing in sample processor 220A. Sample processor 220A for example performing cell separation, cell pre-treatment, pre-concentration or amplification prior to the sample entering a microfluidic chip 220B which is optically interrogated with optical head 220C coupled to opto-electronic circuit 220D. The optical signal is then processed by signal processing electronics 220E which may for example include signal processing, signal encryption, wireless interface, wired interface and logic. The results are presented to the user on display 220F. The body 210A of the POC tester may also include ancillary electronics 220G, such as power supply, USB connector and antenna for example.
Of the many microfluidic systems developed for POC analysis electrokinetically driven microfluidics have existed since the mid-1990 s, see for example “Electrokinetics in Microfluidics, Volume 2 (Interface Science and Technology)” (Elsevier, ISBN-13: 978-0120884445), and whilst powerful require high voltages making for complex systems to operate them. Similarly pneumatically actuated systems have been demonstrated, see for example Braschler et al in “A simple Pneumatic Setup for Driving Microfluidics” (Lab on a Chip, Vol. 7, pp. 420-422), but necessitate large valves. Referring to FIG. 2B there is depicted a chemical reaction driven microfluidic element according to the prior art of Qin et al in “Self-Powered Microfluidic Chips for Multiplexed Protein Assays from Whole Blood” (Lab on a Chip, Vol. 9(14), pp. 2016-2020). Accordingly, a sample is loaded into a chamber within the microfluidic assembly is driven by pressure arising from the generation of oxygen within the microfluidic assembly as a result of a Pt/Ag catalytic breakdown of hydrogen peroxide into water and oxygen which then drives the sample through the microfluidic channels.
Centrifugally driven microfluidics, so called “Lab-on-a-CD” have become popular but require tailor-made spinning systems that have now be made into convenient systems, but are not widely used. Such a Lab-on-a-CD is depicted in FIG. 2C for a five-step flow sequencing CNC-machined CD after Lai et al in “Design of a Compact Disk-like Microfluidic Platform for Enzyme-Linked Immunosorbent Assay” (Anal. Chem., Vol. 76, pp. 1832-1837). Accordingly, capillary microfluidics constitute the most successful technology to date for assays, and indeed are commercially extremely successful for example in providing the lateral flow strips for pregnancy tests, and more recently emergency care for cardiac disease through analysis of a panel of proteins. Such a capillary microfluidic device for immunoassay being depicted in FIG. 2D according to Gervais et al in “Toward One-Step Point-of-Care Immunodiagnostics using Capillary-Driven Microfluidics and PDMS Substrates” (Lab on a Chip, Vol. 9, pp. 3330-3337, hereinafter Gervais2). As depicted a series of functional microfluidic elements are implemented onto the chip for performing immunoassays where the position of and interaction between the analyte, detection antibodies (dAbs) and capture antibodies (cAbs) are illustrated along different parts of the chip where the microfluidic circuit is patterned with lines of cAbs and antigens for the control lines. The silicon circuit has the loading pad, the sample collector, the delay lines, the dAb deposition zones with dAbs, the reaction chamber, the capillary pumps and the vents.
Unlike other microfluidic solutions those based upon microfluidic capillary systems are powered by capillary effects and the control of fluid flow is structurally and chemically encoded into the microscale conduits, the capillaries. According, such capillary systems can be designed to be entirely self-powered and self-regulated, making them very useful for POC applications. Within the prior art an initial library of capillary elements have been reported used for capillary systems including:                microchannels which are closed channels employing hydrophilic or plasma treated conduit surfaces;        serpentine flow resistors which regulate flow rate over a desired region;        delay lines which are typically binary hierarchy microchannels combining flows from a sample collector for example;        vents which are openings within the microfluidic circuit connected to air allowing air to be vented from a closed channel as a fluid fills part of the closed channel;        capillary pumps which are microstructured reservoirs, typically with hydrophilic posts, to generate capillary pressure over a desired region without a significant resistance;        capillary retention valves (CRVs) wherein localized channel cross-section reduction creates a high capillary pressure thereby pinning the fluid after a capillary has been drained; and        capillary trigger valves (CTVs) which are formed as a channel intersects (crosses) a main channel wherein the fluid in the cross-channel is retained for a period of time until its release is triggered by another fluid in the main channel.        
Accordingly, using these elements, capillary systems for filling and draining of one sample at a time can be made, but more complex fluidic operations cannot be achieved with these alone, see for example Juncker et al in “Autonomous Microfluidic Capillary System” (Anal. Chem., Vol. 74, pp. 6139-6144) and Gervais2.
Capillary trigger valves (CTVs), also known as fluidic trigger values, have been proposed by J. Melin et at in “A Liquid-Triggered Liquid Microvalve for On-Chip Flow Control” (12th Int. Conf. Solid State Sensors, Actuators, and Microsystems, 2003, pp. 1562-1565) and Delamarche et al in “Microfluidic Networks for Chemical Patterning of Substrates: Design and Applications to Bioassays” (J. Am. Chem. Soc., Vol. 120, pp. 500-508). However, these are based on high aspect ratios, up to 10 in some cases, of microfluidic channel depth/width and manufactured using deep reactive ion etching (RIE) in silicon (Si). However, such valves are not amenable to fabrication using low cost replica molding techniques that are typically limited to aspect ratios of 2-3. In addition, these valves were prone to leakage owing to flow along the edges at the bottom of the microfluidic channel or have limited retention times. Accordingly it would be beneficial to implement CTVs with low aspect ratio valve that can readily be made using low cost replication techniques, that are robust, do not leak, and have long leadtimes.
Capillary retention valves (CRVs) based upon capillary forces preventing the drainage of fluid have been achieved to date through small cross-section capillary geometries that generates a high capillary pressure (force), see for example Juncker. Initial CRVs were stronger than the capillary pump such that the liquid was retained indefinitely. Whilst developments on CRVs resulted in variable definable retention capacity the retention capacity was not well differentiated and accordingly sometimes CRVs with smaller cross-sections drained before larger cross-sections. Accordingly, it would be beneficial to provide CRVs with reliable retention and release sequence.
The volumes of liquids flushed through capillary systems of microfluidic systems are defined by the geometry of the microfluidic conduits. However, for capillary pumps that are very wide, air bubbles are readily trapped within the capillary pumps because the liquid often proceeds more rapidly along edges, thereby encapsulating a bubble. Zimmerman et at in “Capillary Pumps for Autonomous Capillary Systems” (Lab on a Chip, Vol. 7(1), pp. 119-125, hereinafter Zimmerman1) established a ring based enlargement gradually expanding the lateral dimension of the capillary pump and centering one microstructure in the connecting channel at the entrance of the capillary pump to overcome this. However, filling factors of only approximately 60% to 75% were achieved thereby requiring significantly larger capillary pumps be fabricated to absorb the fluids within the microfluidic circuit. Accordingly it would be beneficial to implement capillary pumps that have increased absorption and avoid the issues of air bubbles.
Further, within prior art microfluidics, such as depicted for example in FIG. 2D the flow within the microfluidic system is one directional. It would be beneficial within some microfluidic systems to not only allow for the autonomous and sequential flow of multiple chemicals at various flow rates but also to provide for flow reversal.
Other aspects and features of the present invention will become apparent to those ordinarily skilled in the art upon review of the following description of specific embodiments of the invention in conjunction with the accompanying figures.